Tissue scaffold having aligned fibrils, apparatus and method for producing the same, and artificial tissue and methods of use thereof

ABSTRACT

A tubular tissue scaffold is described which comprises a tube having a wall, wherein the wall includes biopolymer fibrils that are aligned in a helical pattern around the longitudinal axis of the tube where the pitch of the helical pattern changes with the radial position in the tube wall. The scaffold is capable of directing the morphological pattern of attached and growing cells to form a helical pattern around the tube walls. Additionally, an apparatus for producing such a tubular tissue scaffold is disclosed, the apparatus comprising a biopolymer gel dispersion feed pump that is operably connected to a tube-forming device having an exit port, where the tube-forming device is capable of producing a tube from the gel dispersion while providing an angular shear force across the wall of the tube, and a liquid bath located to receive the tubular tissue scaffold from the tube-forming device. A method for producing the tubular tissue scaffolds is also disclosed. Also, artificial tissue comprising living cells attached to a tubular tissue scaffold as described herein is disclosed. Methods for using the artificial tissue are also disclosed.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a divisional application of co-pending U.S.Non-Provisional patent application Ser. No. 10/861,664, filed Jun. 4,2004, which is related to and claims the priority benefit of U.S.Provisional Application Ser. Nos. 60/475,680; 60/475,866; and60/475,986, filed Jun. 4, 2003, and each of which is incorporated hereinby reference in its entirety.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT

The U.S. Government has a paid-up license in this invention and theright in limited circumstances to require the patent owner to licenseothers on reasonable terms as provided for by the terms of grant no. 1K25 HL67097 awarded by the U.S. National Institutes of Health of theDepartment of Health and Human Services.

BACKGROUND OF THE INVENTION

(1) Field of the Invention

The present invention relates to a tubular tissue scaffold, and, inparticular, to a tubular tissue scaffold having aligned biopolymerfibrils, for use in tissue engineering applications; an apparatus andmethod of producing a tubular tissue scaffold having aligned biopolymerfibrils; and artificial tissue, and methods of use thereof, comprisingliving cells attached to a tubular tissue scaffold having alignedbiopolymer fibrils.

(2) Description of the Related Art

The National Science Foundation defines tissue engineering as “theapplication of principles and methods of engineering and life sciencesto obtain a fundamental understanding of structure-functionrelationships in novel and pathological mammalian tissues and thedevelopment of biological substitutes to restore, maintain, or improve[tissue] function.” See Shalak, R. and Fox, J. eds., Tissue Engineering,Proceedings of a Workshop held at Granlibakken, Lake Tahoe, Calif., Feb.26-29, 1988, New York: Alan Liss (1988). In the last decade, over $3.5billion dollars has been invested worldwide in tissue engineeringresearch. More than 70 start-up companies or businesses having acombined annual expenditure of over $600 million dollars now participatein a significant engineering and scientific effort toward developingalternative sources of transplant materials through in vitro tissueengineering.

Several aspects of creating an engineered tissue make it a dauntingtask. One of the most difficult challenges is directing the behavior ofspecialized cells outside of the body to mimic the normal, endogenousphenotype those cells exhibit in vivo. Additionally, in order for anengineered tissue to be tolerated upon implantation, the material thatprovides the scaffolding for the cells must meet several importantcriteria. The material must be biocompatible, so as not to be toxic orinjurious, and not cause immunological rejection. Also, the materialmust be biodegradable, by having the capability of being broken downinto innocuous products in the body. Because cells respond biologicallyto the substrate on which they adhere, the materials that provide thegrowth surface for engineered tissues must promote cell growth. Further,the scaffolding material should be replaced by extracellular matrixcomponents secreted by the grafted cells as the scaffold is broken downin the body. Additionally, the material should allow cells to grow andfunction as they would in vivo.

Initially, researchers adapted synthetic degradable polyesters that hadbeen used in surgical materials since the early 1970s to constructscaffold materials for use in tissue engineering. These degradablepolyesters include polyglycolide and polylactide, as well as the morerecently developed polymer, polylactide coglycolide. However, thosedegradable polyesters tended to be inflexible, and their degradation invivo has been associated with adverse tissue reactions. Theseshortcomings have led to the development of a host of new syntheticpolymers, for example, polyhydroxybutyrate and copolymers ofhydroxybutyrate with hydroxyvalerate. See Amass, W. et al., Polymer Int,47:89-144 (1998).

In animals, collagens make up a majority of endogenous scaffoldingmaterials. They are the most commonly occurring proteins in the humanbody and they play a central role in the formation of extracellularmatrix. Collagens are triple-helical structural proteins. It is thistriple-helical structure that gives collagens the strength and stabilitythat are central to their physiological role in the structure andsupport of the tissues in the body. Although there are over twenty typesof mammalian collagens, collagen types I, II, III, V, and XI make up thefibrous collagens. Type I collagen molecules polymerize into fibrilswhich closely associate in a parallel fashion to form fibers withenormous tensile strength, which are found in skin, tendon, bone anddentin. Type II is the major collagen found in cartilage, where thefibrils are randomly oriented to impart both stiffness andcompressibility to the proteoglycan matrix. Type III collagen is foundin skin, muscle, and vascular structures, frequently together with typeI collagen.

Collagen has been used successfully in several tissue engineeringapplications. As a copolymer with glycoaminoglycans such as chondroitin6-sulfate, collagen has been utilized as an artificial skin scaffold toinduce regeneration in vivo for the treatment of burn injury since theearly 1980s. See Burke, J. F., et al., Ann Surg, 194:413-28 (1981);Yannis, I. V., et al., Science, 215:174-6 (1982). Collagen has also beenused to form anti-adhesion barriers for use on surgical wounds. See U.S.Pat. Nos. 5,201,745 and 6,391,939 to Tayot et al.; U.S. Pat. No.6,451,032 to Ory et al. Zilla et al., in U.S. Pat. No. 6,554,857,describe the use of collagen, among other materials, as a component in aconcentric multilayer ingrowth matrix that can have a tubular form.

However, some problems remain to be solved in the use of collagen as ascaffolding material for applications requiring structural andmechanical stability, such as for vascular prosthetics. This is at leastpartly due to an inability to isolate collagen possessing the physicalproperties required to maintain necessary mechanical integrity of ascaffold, as it is remodeled in vivo, for use in, for example,cardiovascular indications. Additionally, in order for a tubularconstruct to mimic endogenous components of the cardiovascular system,it must promote the proper growth, orientation, association, andfunction of specialized cell types.

Despite significant work in the field of tissue engineering and thenumerous synthetic biomaterials that have been developed in the lastdecade, there is still a need for improved scaffolding materials for usein specialized applications. It would be useful, therefore, to provide atissue scaffold in the form of a tube, comprising a biopolymer whichpromotes maintenance of an in vivo cell phenotype and, particularly, atubular tissue scaffold that was non-toxic, biologically degradable invivo, and causes little or no immune reaction in a host. It would alsobe useful to provide an apparatus and method for the production of sucha tubular tissue scaffold. Also, despite the advances in biomaterialsresearch, and the elucidation of the molecular biology of cell behaviorand cell:matrix interactions, the gap between in vitro engineered tissueand biologically functional implantable organs remains significant.Therefore, it would be useful to provide artificial tissue that can actas a functional prosthetic. It would also be useful if the artificialtissue could be produced in the form of a tube, utilizing the tissuescaffolding described herein. This structural configuration would beparticularly useful in cardiovascular applications.

SUMMARY OF THE INVENTION

Briefly, therefore, the present invention is directed to a novel tubulartissue scaffold comprising a tube having a wall, wherein the wallincludes biopolymer fibrils that are aligned in, a helical patternaround the longitudinal axis of the tube where the pitch of the helicalpattern changes with the radial position in the tube wall.

The present invention is also directed to a novel apparatus forproducing a tubular tissue scaffold having aligned biopolymer fibrils,the apparatus comprising a biopolymer gel dispersion feed pump that isoperably connected to a tube-forming device having an exit port, wherethe tube-forming device is capable of producing a tube from the geldispersion while providing an angular shear force across the wall of thetube, and a liquid bath located to receive the tubular tissue scaffoldfrom the tube-forming device.

The present invention is also directed to a novel method of producing atubular tissue scaffold, the method comprising:

providing a gel dispersion comprising a biopolymer;

feeding the gel dispersion to a tube-forming device that is capable ofproducing a tube from the gel dispersion while providing a radial shearforce across the wall of the tube and having a gas channel connecting agas source with the luminal space of the tubular tissue scaffold as itexits the tube-forming device;

forming the gel dispersion into a tube; and

causing the gel dispersion to solidify, thereby forming a tubular tissuescaffold comprising a tube wall having biopolymer fibrils that arealigned in a helical pattern around the longitudinal axis of the tubeand where the pitch of the helical pattern changes with the radialposition in the tube wall.

The present invention is also directed to novel artificial tissuecomprising living cells attached to a tubular tissue scaffold comprisinga tube having a wall, wherein the wall includes biopolymer fibrils thatare aligned in a helical pattern around the longitudinal axis of thetube where the pitch of the helical pattern changes with the radialposition in the tube wall.

The present invention is also directed to a novel method ofpreconditioning artificial tissue for implantation into the body of asubject, the method comprising: seeding a tubular artificial tissuescaffold having aligned biofibrils with living cells and culturing thecells in the presence of media containing at least one growth factor andunder conditions where the tubular artificial tissue scaffold issubjected to stretch and pressure pulse of controlled frequency andamplitude.

The present invention also includes a novel preconditioned artificialtissue comprising living cells that are attached to a tubular artificialtissue scaffold having aligned biofibrils, wherein the cells have beencultured in the presence of media containing at least one growth factorand under conditions where the tubular artificial tissue scaffold issubjected to stretch and pressure pulse of controlled frequency andamplitude.

Additionally, the present invention provides a method of treatmentcomprising implanting in the body of a subject artificial tissuecomprising living cells attached to a tubular tissue scaffold comprisinga tube having a wall, wherein the wall includes biopolymer fibrils thatare aligned in a helical pattern around the longitudinal axis of thetube where the pitch of the helical pattern changes with the radialposition in the tube wall.

The present invention also encompasses a method of identifying theeffects of a pharmaceutical composition on cell function comprisingadministering said pharmaceutical composition in vitro to artificialtissue comprising living cells attached to a tubular tissue scaffoldcomprising a tube having a wall, wherein the wall includes biopolymerfibrils that are aligned in a helical pattern around the longitudinalaxis of the tube where the pitch of the helical pattern changes with theradial position in the tube wall.

Among the several advantages found to be achieved by the presentinvention, therefore, may be noted the provision of a tubular tissuescaffold having sufficient structural strength to withstand pressure;the provision of a scaffold having a composition that is biologicallydegradable in vivo and will result in a minimal immunological responsefrom a host; the provision of a tissue scaffold having a structuralcomposition that allows the penetration of cells and provides cells therequisite signals to develop an in vivo functional phenotype; theprovision of an apparatus and a method for the production of such atubular tissue scaffold, and; the provision of artificial tissuecomprising cells attached to a tissue scaffold of biopolymer fibrilsthat is non-immunogenic, has a construction that mimics that of cardiactissue, and has improved structural integrity.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is an illustration showing a partial cross-sectional cutaway viewof a tubular tissue scaffold of the present invention illustratinghelical alignment patterns of different direction of biopolymer fibrilsat the luminal surface (left-hand pitch) and outside surface (right-handpitch) of the tube (A), and also showing three side views of tubulartissue scaffolds showing that the pitch of the helical alignment patternof fibrils at the luminal surface (dotted line having an angle with thelongitudinal axis of α_(L)) and outside surface (solid lines having anangle with the longitudinal axis of α_(O)) of the tube can beapproximately equal, but opposite in direction (B), or unequal in pitch,but opposite in direction (C), and include the case where the biopolymerfibrils are aligned at zero pitch (α_(O)=0°), at the outside wall inthis case, while the biopolymer fibrils at another location in the tubewall (here, the luminal surface of the tube wall) are aligned in aright-hand pitch of angle α_(L) (D); and

FIG. 2 shows scanning electron micrographs of the luminal and outsidesurfaces of a tubular scaffold of the present invention in which thelongitudinal axis (along the center of the tube) is roughly verticalbetween the two images and where panel (A) shows the alignment of thecollagen fibrils on the outside of the tube wall and panel (B) shows thereverse pitch orientation of the collagen fibrils on the luminal surfaceof the tube wall;

FIG. 3 is a sectional view illustrating the major parts of a counterrotating cone extruder of the present invention;

FIG. 4 is a schematic illustration of the major parts of an apparatusfor producing a tubular tissue scaffold of the present invention;

FIG. 5 shows a laser scanning confocal micrograph (A) of cardiacmyocytes on a collagen scaffold of the present invention, stained withantibodies to F-actin and connexin 43, illustrating expression ofconnexin 43 along the sides and ends of the myocytes. Panels (B) and (C)are a stereo pair of laser scanning confocal micrographs of cardiacmyocytes on a collagen scaffold of the present invention stained with anantibody to F-actin (Alexa 488 phalloidin). Multiple layers of cellswith aligned fibrils can be seen when viewing the images with stereoviewing equipment;

FIG. 6 shows transmission electron micrographs of cardiac myocytes on acollagen scaffold of the present invention. The micrographs demonstratethe presence of Z-bands (Z), aligned microfibrils (F), numerousmitochondria (M), cell:cell junctions (arrows in inset A), andinteraction with collagen in the tube wall (arrows in inset B);

FIG. 7 shows a representative electrical signal obtained from a tubularscaffold of the present invention seeded with neonatal cardiac myocytesin panel (A), and Fast Fourier transform analysis of the electrical datain panel (B); and

FIG. 8 shows a schematic illustration of a system including a pulsatileinternal flow stretch bioreactor wherein the tubular tissue scaffold(tissue engineered tube) is attached at each end to supply and effluenttubing and media is fed through the tubing. A basal level of pressure isprovided due to the height of the media reservoir and pulsed flow tocyclically stretch the cells. The bioreactor stretches the tubularscaffold both radially and longitudinally.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

In accordance with the present invention, it has been found that atubular tissue scaffold can be produced that comprises a tube wall whichincludes biopolymer fibrils that are aligned in a helical pattern aroundthe longitudinal axis of the tube where the pitch of the helical patternchanges with the radial position in the tube wall, using a novelapparatus and method of producing such a tubular tissue scaffold. It hasalso been found that artificial tissue comprising living cells attachedto the novel tubular tissue scaffold described herein can be producedand used to treat a variety of disease conditions by implanting theartificial tissue in the body of a subject.

The novel scaffold is unique in that it is not a very soft gel withlittle mechanical integrity, nor is it a tough highly crosslinkedmaterial resembling leather. In addition, toxic or potentially cytotoxicchemicals are not required to crosslink the scaffold material. Rather,the novel scaffold is made as an extruded tube with aligned biopolymerfibrils that impart biochemical and biomechanical information to cellsthat attach to them, which instruct the cells to adopt an in vivo-likephenotype.

The novel tissue scaffold provides the advantages that it has sufficientstructural strength to maintain its tubular conformation without asupport such as a mandrel, and can withstand at least about 12 mmHginternal pressure and in preferred forms can withstand pressures of atleast about 280 mmHg, or more. The scaffold has a composition that willresult in a minimal immunological response from a host; and the scaffoldhas a structural composition that allows penetration by living cells andprovides cells the requisite signals to develop an in vivo functionalphenotype.

From a mechanical standpoint, the properties provided by biopolymersthat form fibrils give the scaffold a suitable modulus, suitableflexural rigidity, and a surface that is not too hydrophobic or toohydrophilic to support the cells and fluid necessary to form a tissue.Moreover, the biopolymers are biodegradable, nontoxic and support themaintenance and synthesis of new tissue. Furthermore, the presentscaffold material will biodegrade at a controlled rate to allow newtissue to invade.

Pore size and distribution have been reported as important parametersfor scaffold characterization and efficacy. See Dagalakis, N. et al., J.Bio Med Matl Res. 14(4):511 (1980), and Zeltinger, J. S. et al, TissueEngineering. 7(5):557 (2001). The present tissue scaffold material isporous, with most pores between 1 and 10 μm in diameter. Although theaverage size of cells such as neonatal rat cardiac myocytes, forexample, is 15-18 μm, it is believed that pore size is not a limitingfactor that prevents myocyte penetration into the present scaffoldmaterial, since cells can remodel and create an appropriateextracellular matrix material if placed in the right environment. Poresize in the novel tissue scaffold material is discussed further below.

The present tubular tissue scaffolds provide biopolymer fibrils that arealigned in the tube wall in a helical fashion around the longitudinalaxis of the tube. It is important, however, that in the presentscaffolds the pitch of the helical pattern of the aligned fibrilschanges as the radial location in the tube wall progresses from theinside (luminal) surface of the tube wall to the outside (or, outer)surface of the tube wall. In some embodiments, this pattern of fibrilalignment can be characterized as being in a “corkscrew” pattern at theluminal wall of the tube and changing to a “counter-corkscrew” patternat the outside wall of the tube. In other words, if fibril alignment onthe luminal surface is in a right-hand helical pitch, the fibrilalignment on the outside surface of the tube would be in a left-handpitch, or vice-versa. This pattern mimics the extracellular matrixpattern of heart tissue, and, in one embodiment, allow the constructionof more biologically-similar vascular constructs and heart constructs.By way of example, when the tissue scaffold is used to support thegrowth of neonatal cardiac myocytes, the myocytes can be made toconstrict simultaneously, thereby twisting the tube while reducing itsdiameter, and thereby “wringing” fluid from the tube in a motion similarto that of a biological heart.

The present tubular tissue scaffold differs from previous constructs,such as those described in U.S. Pat. No. 6,540,780 to Zilla et al.,which describes synthetic vascular grafts with helically orientedingrowth channels within the tube wall, and U.S. Pat. No. 6,554,857 toZilla et al., in that the present tubular tissue scaffolds comprisebiopolymer fibrils that are aligned in a helical pattern within the tubewall. In Zilla et al. '780, only synthetic polymers are used, and bothof these patents describe constructs where it is the ingrowth channelsthat are aligned within the tube wall, rather than biopolymer fibrils.In order to obtain the present tubular scaffolds, it is necessary toform the tube wall in a manner that imposes an angular shear across thetube wall as it is being formed, as is described below. The presenttubular scaffolds, in fact, are substantially free of helically orientedchannels within the tube walls.

When discussing the geometry of a tubular scaffold herein, the axisalong the center of a cylindrical tube will be referred to as thelongitudinal axis. The axis that is perpendicular to the longtudinalaxis and runs outward from the center of the tube in a direction that isperpendicular to the tube wall is referred to as the radial axis. Thehelical pattern of the aligned biopolymer fibrils within the tube wallis described in terms of the “pitch” of the helix. As in the descriptionof screw threads, for example, the pitch of the helix can be “righthand” or “left hand”. The pitch of the helical pattern formed by thealigned fibrils is described in terms of the angle (α) between a tangentto the helix and a projection of the longitudinal axis of the tube (asshown, for example, in FIG. 1, of the present specification).

The pitch of the helical pattern of the biopolymer fibers at the luminal(internal) wall of the tube is designated as αL and the pitch of thehelical pattern of the biopolymer fibers at the outside wall of the tubeis designated as α_(O).

When it is said that the biopolymer fibrils are “aligned”, it is meantthat most of the fibrils in the same radial plane in a tube wall runroughly parallel to each other. It is not meant that every fibril mustbe parallel to every other fibril in the plane, but that a generalalignment pattern must be discernable. Such fibril alignment is shown,for example, in FIG. 2. The fibrils of the present tubular tissuescaffold are preferably not isolated from each other in the wall of thetube, but, rather, are associated in sheets or areas containing dozens,if not hundreds or thousands of fibrils that are adjacent or touchingand are associated and a part of the helical pattern of alignment.

The term “fibrils” is used herein to describe an association of severalbiopolymer molecules into a structure that appears fibrous with suitablemagnification, as is typical for collagen and other selected biopolymersof biological origin. There is no particular size limit to the fibrilsof the present invention, and both fibrils and small fibers are includedin the term.

As used herein, the term “biopolymer” refers to a natural or syntheticpolymer that is biologically compatible. Polymers that are biologicallycompatible are those which can be implanted into a living vertebratesubject without triggering a severe adverse immune reaction. Examples ofbiopolymers that can be used in the present invention include, but arenot limited to, collagen, fibronectin, laminin, elastin, fibrin,proteoglycans, hyaluronan, and combinations thereof. In someembodiments, the biopolymer is a collagen, selected from the groupconsisting of type I collagen, type II collagen, type III collagen, typeV collagen, type XI collagen, and combinations thereof. In oneembodiment, the collagen is selected from the group consisting of type Icollagen, type III collagen, and combinations thereof. In preferredembodiments, the biopolymer is type I collagen.

Biopolymers that have not been isolated or purified to some degree fromtheir natural sources, however, are not included in the scope of thepresent invention. In other words, the present tubular tissue scaffoldis not meant to include natural vessels, arteries, or other naturalbiological tubular structure.

Type I collagen is the most prevalent structural extracellular matrixprotein in the human body. Collagen has the property of being able todirect cell behavior by signaling the cells to modify their growth,differentiation, intercellular contacts, and production of moleculessuch as collagen and other extracellular matrix proteins and cytokines.Additionally, cells are capable of recognizing and correctly modifyingthe collagen matrix to conform to their cellular requirements. Thesefeatures allow cells introduced to a collagen scaffold to mimic theirnormal in vivo phenotype and organization. By way of example, cells suchas myocytes introduced to the surfaces of the tubular scaffold wouldreadily attach to the collagen, remodel the matrix, and developintercellular connections.

The tubular tissue scaffold of the present invention comprises a tubehaving defined dimensions such as an outside diameter, a luminaldiameter, and a wall thickness. The novel tubular tissue scaffold canhave any dimensions and is not limited to any particular diameter orwall thickness.

In one embodiment, the outside diameter of the tubular tissue scaffoldis between about 0.1 millimeter and about 100 millimeters. In someembodiments, the outside diameter is between about 0.5 millimeters andabout 50 millimeters, or between about 1 and about 10 millimeters, orbetween about 4 and about 10 millimeters. In preferred embodiments, theoutside diameter is about 5 millimeters.

The luminal diameter of the tubular tissue scaffold is smaller than theoutside diameter of the tube and may be between about 0.1 millimetersand about 49 millimeters. In one embodiment, the luminal diameter isbetween about 0.4 millimeters and about 49 millimeters. In preferredembodiments, the luminal diameter is between about 4 millimeters andabout 5 millimeters.

The thickness of the tube wall, along with the size of the lumen,determines the rate of diffusion of nutrients that are critical for cellgrowth from the outside of the tube to the luminal surface. The tubewall can be of any thickness that will provide properties that aresuitable for the intended application. In one embodiment, the tube wallthickness is between about 0.05 millimeters and about 10 millimeters. Inanother embodiment, the wall thickness is between about 0.1 millimetersand about 5 millimeters. In preferred embodiments, the wall thickness isbetween about 0.1 millimeters and about 1 millimeter.

As has been described briefly above, the tube wall of the tubular tissuescaffold comprises a helical alignment of biopolymer fibrils, thealignment having a certain pitch in relation to the longitudinal axis ofthe tube. A purpose of the particular alignment of biopolymer fibrils ofthe present invention is to allow for the generation of a pumping, or“wringing” type of mechanical force in a tubular scaffold construct,given the introduction of the proper cells to the scaffold.

In the present invention, the pitch of the helical pattern of biopolymerfibrils on the luminal surface can be between about 0 degrees and about89 degrees and the pitch of the helical pattern on the outer surface isbetween about 0 degrees and about 89 degrees, where the pitch of thefibrils at the luminal surface is different from the pitch of thefibrils at the outer surface of the tube. In one embodiment, the pitchof the helical pattern of the fibrils at the luminal surface is betweenabout 18 degrees and about 62 degrees and the pitch of the helicalpattern of the fibrils at the outer surface is between about 18 degreesand about 62 degrees, where the pitch of the fibrils at the luminalsurface is different than the pitch of the helical pattern of thebiopolymer fibrils at the outer surface. In preferred embodiments, thepitch of the helical pattern of the biopolymer fibrils at the luminalsurface is between about 26 degrees and about 60 degrees and the pitchof the helical pattern of the biopolymer fibrils at the outer surface ofthe tube is between about 26 degrees and about 60 degrees, where thepitch of the fibrils at the luminal surface is different from the pitchof the fibrils at the outer surface. In another embodiment, the pitch ofthe helical pattern of the biopolymer fibrils changes in a linear mannerthrough the thickness of the tube wall, and the change can be from aright-hand pitch to a left-hand pitch, or vice versa.

In one embodiment of the present invention, the tubular tissue scaffoldhas pores. One feature of pore size is the effect that it has on theability of cells to invade the matrix. In preferred embodiments of thepresent invention in which a biopolymer, such as collagen, for example,is used, cells introduced to the scaffold are able to remodel thematerial as they would in vivo, enabling the cells to infiltrate thematrix. Pore size must also provide sufficient permeability for thediffusion of nutrients from the media in in vitro culture conditions. Inpreferred embodiments of the present invention, the size of the pores isfrom about 1 micron to about 20 microns. More preferably, the size ofthe pores is from about 2 microns to about 10 microns. Unless explicitlydescribed to the contrary herein, pore size is expressed in terms ofnumber average pore size.

In some embodiments of the tubular tissue scaffold of the presentinvention, the luminal surface and outer surface of the tube can supportcell attachment and growth. In preferred embodiments, the tissuescaffold can direct the morphology of attached cells to align along thehelical pattern of the biopolymer.

In addition to growth and attachment of cells at the luminal surfaceand/or the outer surface of the tube, it is sometimes desirable thatgrowing cells penetrate and grow into the tube wall. In other words,that growing cells grow throughout the depth or thickness of the tissuescaffold material. An advantage of the present tissue scaffold is thatcells are able to penetrate and grow throughout the entire thickness ofthe tube wall, and in preferred embodiments, the growing cells alignwith the general alignment of the biopolymer fibrils throughout thethickness of the wall. This provides for a construct of artificialtissue scaffold having changing alignment of biopolymer fibrilsthroughout the tube wall thickness and also having growing cellsthroughout the tube wall thickness that match the alignment of thebiopolymer fibrils. This structure permits closer matching of thephysical structure of the artificial tissue with that of normalcardiovascular tissue.

In one embodiment, the tissue scaffold of the present invention can betreated with UV radiation to crosslink the biopolymer fibrils making upthe tissue scaffold to increase the strength of the scaffold, asmeasured by increased burst pressure. As used herein, the term “burstpressure” refers to the maximum amount of fluid pressure which may beapplied to the tubular tissue scaffold internally without causing arupture. Treatment with radiation will be further described below, butthe inventors have found that treatment of tubular tissue scaffolds withUV radiation increases the strength of the artificial tissue scaffold.When the artificial tissue scaffold is in the form of a tube, itsstrength can easily be measured and expressed in terms of the “burstpressure”. As used herein, the terms burst pressure mean the internalpressure at which a tube bursts. Treatment of a present tubularartificial tissue scaffold with UV radiation at a wavelength that isfrom about 250 to about 280 nm and an energy density of from about 100to about 1000 μw/cm², is preferred. By way of example, treatment with UVradiation at a wavelength of 254 nm and an energy density of 500 μw/cm²for 140 minutes increased the burst pressure of the tubes to 280 mm Hg,as compared to untreated tissue scaffolds which have an average burstpressure of about 125 mm Hg.

In preferred embodiments, a tubular tissue scaffold of the presentinvention has a burst pressure of at least about 100 mm Hg, morepreferably at least about 200 mm Hg, even more preferably at least about250 mm Hg, and most preferably at least about 280 mm Hg.

In another embodiment, the present invention is directed to an apparatusfor use in the production of a tubular tissue scaffold having alignedbiopolymer fibrils, as described herein.

The apparatus of the present invention can be described with referenceto the figures. In FIG. 4 an apparatus for producing a tubular tissuescaffold having aligned biopolymer fibrils is illustrated. The apparatuscomprises a gas source (200), which can be any type of source that cansupply gas having a desired composition. It is preferred that the gas isa mixture of air and ammonia, and a mixture of air and ammonia in abouta 50:50 mixture by volume is more preferred. A compressed gas cylinderis commonly used as a gas source, but the apparatus is not limited tosuch a source. The gas source (200) is operably connected to provide gasto a tube-forming device (100) through a conduit (201), and can, ifdesired, be operably connected to provide gas to a controlled atmospherechamber (600) through a conduit (202). The gas source (200) can be usedto supply gas to the luminal space of the tubular tissue scaffold (700),and, if desired, to the interior of the controlled atmosphere chamber(610). In a preferred embodiment, the controlled atmosphere chamber isfilled with a gas mixture comprising air and ammonia, and a gas mixturecomprising a mixture of about 50:50 air and ammonia by volume is morepreferred.

A biopolymer feed pump (300) is operably connected to feed a biopolymergel from a biopolymer gel source to the tube-forming device (100). Thefeed pump can be any type of liquid feeding device. Examples of suitablepumps are centrifugal pumps, gear pumps, peristaltic pumps, progressingcavity pumps, syringe pumps, and the like. It is preferred that the pumpis one that is capable of feeding a viscous biopolymer gel (perhapshaving a viscosity of from about 10 centipoise to as much as 1000centipoise, or even more), at a metered rate, and under conditions wherethe cleanliness, and even sterility, of the biopolymer gel can bemaintained. A syringe pump has been found to be useful for small scaletube-forming systems.

The tube-forming device (100) is one that is capable producing a tubefrom the gel dispersion while providing an angular shear force acrossthe gel dispersion as it is formed into the wall of the tube. Thetube-forming device (100) is preferably located so that the tubulartissue scaffold exiting the device (700) can easily be transferred to aliquid bath (500) located to receive the tubular tissue scaffold fromthe tube-forming device. The liquid bath contains a liquid that receivesand cushions the extruded tube. The bath also serves to facilitate thesolidification of the tube. In an embodiment, the liquid bath compriseswater having sufficient ammonia absorbed therein to bring the pH of thewater to between about 9 and 11, a pH of about 10 is preferred.

It is also preferred that the space between the exit of the tube-formingdevice (100) and the liquid bath (500) be enclosed to provide acontrolled atmosphere chamber (600). In embodiments where the tubulartissue scaffold descends from the exit of the tube-forming device (100)to the liquid bath (500) by force of gravity, it is desirable that thesurface of the liquid in the bath (510) be located a defined distance(L) below the exit of the tube-forming device.

In a preferred embodiment, the surface of the liquid bath is locatedbetween about 0.25 centimeter and about 60 centimeters below the exitport of the extruder, a distance of between about 2.5 centimeters andabout 25 centimeters is more preferred, and a distance of between about7.6 centimeters and about 16.2 centimeters is even more preferred.

In a preferred embodiment, the tube-forming device (100) is a rotatingcone extruder or a rotating disk extruder. In a yet more preferredembodiment, the tube-forming device is a rotating cone extruder. In aneven more preferred embodiment, the tube-forming device is acounter-rotating cone extruder. With reference to FIG. 3, the majorparts of a counter-rotating cone extruder include a body (105) withcover plate (104), within which is contained an external rotating member(120) having a cone-shaped cavity therethrough and an internal rotatingcone (110), which may also be referred to herein as an “internal member”or internal rotating member', and which fits within the cone-shapedcavity of the external rotating member. The internal cone-shaped member(110) is connected to and is driven through an internal member drivegear (112) and the external rotating member (120) is connected to anddriven by an external member drive gear (122). Both the internal memberdrive gear (112) and the external member drive gear (122) can engage andbe driven by a pinion gear (130) that is connected to and driven by adrive shaft from the drive motor (400). The external rotating member andthe internal rotating cone can terminate near the apex of the cones toform an annular-shaped exit port (150).

The external rotating member (120) and the internal rotating cone (110)can be operably connected to one or more drive motors (400) that canspin the external rotating member and the internal rotating cone about acommon axis but in opposite directions. In a preferred embodiment, theexternal rotating member and the internal rotating cone are operablyconnected to the same drive motor. In a preferred embodiment, the one ormore drive motors can be adjusted to vary the speed at which theexternal rotating member and the internal rotating cone rotate. Thedrive motor(s) can be designed to rotate at any speed, but it ispreferred that the motor(s) be designed so that the rotational speed ofthe external rotating member and the internal rotating cone can bevaried between about 1 rpm and 1800 rpm.

The counter-rotating cone extruder of the present invention can besmaller than counter-rotating cone extruders that are in common use andare commercially available. One of the problems associated with themanufacture of a reasonably priced counter-rotating cone extruder is theprovision of the intricate bearings needed to retain the alignment andspacing of the external rotating member and the internal rotatingmember. It has been found that bearings can be dispensed with if theexternal rotating member and the internal rotating cone of the extruderare constructed of a durable, bearing-quality polymer. In preferredembodiments, the external rotating member and the internal rotating coneare constructed of Delrin® acetal resin (available from Tri Starplastics, Reading, Pa.).

If desired, the annular exit port (150) of the extruder isinterconnected with a gas source via a gas conduit (140) to provide forthe addition of gas to the luminal space of a tubular tissue scaffoldexiting the extruder.

When the tube-forming device is a rotating-cone extruder or acounter-rotating cone extruder, the annular-shaped exit port has anoutside diameter and an inside diameter. The difference between theoutside diameter and the inside diameter of the annular space definesthe width of the annular space. The thickness of the wall of the tubulartissue scaffold is directly related to, but not equivalent to, the widthof the annular space.

The outside diameter and the inside diameter of the annular space can beany desired dimension, but it is preferred that the outside diameter ofthe annular-shaped exit port is between about 0.5 mm and about 150 mm,and the width of the annular space is between about 0.1 mm and about 10mm. It is more preferred that the outside diameter of the annular-shapedexit port is between about 1 mm and about 20 mm, and the width of theannular space is between about 0.1 mm and 2 mm. It is yet more preferredthat the outside diameter of the annular-shaped exit port is betweenabout 1 mm and about 10 mm, and the width of the annular space isbetween about 0.1 and 1 mm.

The present invention also includes a method of producing a tubulartissue scaffold, the method includes the provision of a gel dispersioncomprising a biopolymer.

In addition to the isolated and purified biopolymer, or combination ofbiopolymers, the biopolymer fibrils that form the tube wall of thepresent tissue scaffold can be mixed with any other polymers, or otheradditives, that are useful for the formation of, or the performance of,the tubular tissue scaffold. For example, fillers, dyes, drugs, or anyother useful and pharmacologically acceptable material may be added.

The biopolymer is prepared to form a gel dispersion prior to formationof the present tubular tissue scaffold. By way of example, a geldispersion of type I collagen can be prepared by: a) washing bovine hidesequentially in water, water containing NaHCO₃ and surfactant, andwater; b) contacting the hide with an aqueous solution containingNaHCO₃, Ca(OH)₂ and NaHS; c) washing the hide in water; d) treating thehide with an aqueous solution of Ca(OH)₂; e) rinsing the hide with waterand trimming the hide of any remaining skin tissue and fat; f) placingthe hide in an aqueous salt solution and adding hydrochloric acidsolution until the pH is stable between about 6.0 and 8.0; g) washingthe hide in water; h) placing the hide in an aqueous solution of aceticacid with or without pepsin; i) mixing and allowing the hide to swell;j) placing the swollen hide in a mill and processing into a geldispersion; k) filtering the gel dispersion to remove undissolvedparticles; l) centrifuging the gel dispersion to remove smallundissolved particles; m) adding salt to the gel dispersion in an amountsufficient to precipitate collagen from the gel dispersion; n) filteringthe collagen precipitate and resuspending it in deionized water; o)adding a base to bring the pH of the collagen dispersion to a pH betweenabout 6 and about 8; p) dialyzing the collagen dispersion againstphosphate buffered saline solution or tris[hydroxymethyl]aminomethanebuffer; q) resuspending the collagen in deionized water; r) centrifugingthe collagen dispersion to concentrate solid collagen gel as a pellet;and s) resuspending the pellet in aqueous mineral acid or organic acid.When an organic acid is used, acetic acid is preferred. When the pelletis resuspended in acid, it is preferred that the collagen concentrationis adjusted to between about 15-35 g/l, about 20 g/l is more preferred.

The concentration of the gel dispersion can be adjusted to contain about2%-3%, by weight, solids by the addition of water. The gel dispersioncan be fed to a tube-forming device that is capable producing a tubefrom the gel dispersion while providing an angular shear force acrossthe wall of the tube. This device can be a counter-rotating coneextruder, a counter-rotating disk extruder, or a counter-rotatingcylinder extruder that preferably has a gas channel connecting a gassource with the luminal space of the tubular tissue scaffold as it exitsthe tube-forming device.

When the terms “angular shear force” are used herein, what is meant is ashear force that is applied across the wall of the tube, and in adirection generally perpendicular to both the radial and longitudinaltubular dimensions. In other words, a shearing force from the luminalwall of the tube to the outer wall of the tube, or vice versa, in acircumferential direction—as provided, for example, be a rotating coneextruder.

The tube-forming device forms the gel dispersion into a tube, and thetube is then solidified, thereby forming a tubular tissue scaffoldcomprising a biopolymer having fibrils in the tube wall that are alignedin a helical pattern around the longitudinal axis of the tube and wherethe pitch of the helical pattern changes with the radial position in thetube wall. Often, the pitch of the helical pattern on the luminalsurface of the tube is different from the pitch of the helical patternon the outside surface of the tube.

The pitch of the helical pattern of the fibrils can be controlled bycontrol of such variables as the feed rate of the biopolymer geldispersion to the tube-forming device, the rate of shear imposed on thegel as the tube is being formed, and the degree of drawing orcompression of the tube after formation, but before solidification. Insome embodiments, where the tube-forming device is a counter-rotatingcone extruder that has an external rotating member having a cone-shapedcavity therethrough and an internal rotating cone which fits within thecone-shaped cavity of the external rotating member, the externalrotating member is driven to rotate in one direction at a speed of fromabout 1 to about 1800 rpm and the internal rotating cone is driven torotate in the opposite direction at a speed of from about 1 to about1800 rpm. In a preferred embodiment, the external rotating member isdriven to rotate in one direction at a speed of from about 150 to about900 rpm and the internal rotating cone is driven to rotate in theopposite direction at a speed of from about 150 to about 900 rpm.

In a preferred embodiment of the present apparatus, the biopolymer feedpump (300) can be adjusted to vary the rate at which biopolymer is fedto the tube-forming device. When the tube-forming device is an extruderand the liquid bath (500) is located at a defined distance (L) below theextruder exit port (150), the tubular tissue scaffold exiting theextruder (700) can fall into the bath by force of gravity. In thisconfiguration, it is possible to control the conformation of the tube asit solidifies by controlling the biopolymer feed rate and the distancebetween the exit port of the extruder and the surface of the liquid inthe bath (L). Because the tube exiting the extruder is still anunsolidified gel dispersion, the tube can be drawn orcompressed—affecting the aligned helical pattern of the fibrils—or itcan collapse, unless certain measures are taken to control theconformation of the unsolidified tube.

In order to prevent the tube from collapsing, a gas can be fed from thegas source to the luminal space of the tubular tissue scaffold as itexits the extruder. The flow rate of this gas can be controlled so thatit provides an internal pressure inside the tube sufficient to preventthe collapse of the walls of the tube without causing undue expansion ofthe tube that would adversely affect the helical pattern of the alignedfibrils. When collagen is used as the biopolymer, it is preferred thatthe gas comprises a mixture of air and ammonia gas. Contact of theammonia with the biopolymer dispersion causes the pH of the walls of thetube to raise quickly, thereby facilitating the solidification of thebiopolymer gel. In preferred embodiments, the mixture of air and ammoniais about a 50:50 mixture by volume.

In order to further facilitate the solidification of the tube ofbiopolymer gel, the controlled atmosphere chamber can be filled with thesame gas as is fed to the lumen of the tube and the outside surface ofthe tubular tissue scaffold is contacted with a mixture of air andammonia gas as it exits the extruder. Solidification of the tubulartissue scaffold can also be facilitated by providing that the liquid ofthe liquid bath is composed of water having sufficient ammonia dissolvedtherein to raise the pH of the bath liquid to about 10.

As mentioned above, when the liquid bath is located beneath the exitport of the extruder, the conformation of the tube can be controlled asit solidifies by controlling the biopolymer feed rate and the distancebetween the exit port of the extruder and the surface of the liquid inthe bath (L). This can be accomplished by feeding the biopolymer geldispersion to the extruder at a defined feed rate, and the defined feedrate and the defined distance (L) are selected so that the geldispersion solidifies to form a tubular tissue scaffold comprising abiopolymer having fibrils in the tube wall that are aligned in a helicalpattern around the longitudinal axis of the tube and where the pitch ofthe helical pattern on the luminal surface of the tube is different fromthe pitch of the helical pattern on the outside surface of the tube.

In an embodiment where the outside diameter of the annular-shaped exitport is about 5 mm and the width of the annular space is about 0.5 mm,the defined feed rate can be controlled to be sufficient to provide atube extrusion rate of about 150 cm/min. and the defined distance (L) isbetween about 10 cm and about 20 cm. The “tube extrusion rate” referredto is the linear rate of speed at which the tubular tissue scaffoldexits the tube-forming device.

Although the tubular tissue scaffold can remain in the bath for anydesired length of time, when the bath comprises aqueous ammonia having apH of about 10, it is preferred to leave the tissue scaffold in the bathfor about 10 min-20 min, and about 15 min. is more preferred.

In a preferred embodiment, the step of causing the gel to solidify canfurther include immersing the tube in an aqueous solution containing0.3% sodium bicarbonate, percent by weight.

After the gel has solidified, the tube can be sterilized by exposure tosterilizing radiation. Gamma radiation and/or UV radiation can be usedfor this purpose. In some embodiments, the use of both gamma and UVradiation is preferred.

The tubular tissue scaffold of the present invention could be used, forexample, as an implanted prosthesis. In and of itself, the scaffoldcould act as a growth substrate on which the cells of the subject inwhich it is implanted adhere, replicate, and reconstruct the injuredtissue. Additionally, the tubular tissue scaffold could be seeded withspecific cell types in vitro, cultured, and then implanted in a subject.

In another embodiment of the present invention, the tubular tissuescaffold is split along the longitudinal axis and opened to form asheet. The sheet contains layers of aligned biopolymer fibrils, wherethe direction of the alignment changes in each successive layer. As withthe tubular construct, the sheet could be seeded in vitro with specificcell types either before or after splitting. Applications for this typeof tissue scaffold could include any application that requires asheet-type of tissue, rather than a tubular structure. Sheet-typescaffold material can be used, for example, for the preparation ofartificial skin to treat burn injury or surgical patches for internalapplication.

In another embodiment, the present invention encompasses artificialtissue comprising living cells attached to a tubular tissue scaffoldcomprising a tube having a wall, wherein the wall includes biopolymerfibrils that are aligned in a helical pattern around the longitudinalaxis of the tube where the pitch of the helical pattern changes with theradial position in the tube wall.

In an alternative embodiment, when the tubular artificial tissuescaffold is split to form a sheet-type structure, the artificial tissuecomprises a sheet having a thickness, wherein the material comprisingthe sheet includes biopolymer fibrils having an alignment that changeswith the thickness of the sheet.

In the artificial tissue of the present invention, living cells may beintroduced to the luminal and outer surfaces of the tubular tissuescaffold. As used herein, the terms “introduced”, “seed”, and “seeded”in reference to cells refer to the addition of cells to the tissuescaffold by providing the cells in a cellular suspension andsupplementing the solution in which the tissue scaffold is incubatingwith the cellular suspension. By way of example, cells are isolated froma given source, such as neonatal rat hearts, and dispersed in a solutionto form a cellular suspension having a certain cell density. A volume ofcellular suspension is injected into the tubular tissue scaffold usingan IV catheter. The tubes are placed in a rotating wall bioreactor andthe reactor is filled with additional cell suspension. The tubes andcells are incubated at a rotation rate of 20 rpm with 5% CO₂ at 37degrees Celsius for several days.

In some embodiments, the living cells align along the helical pattern ofthe biopolymer. “Aligned along the helical pattern of the biopolymer”,as used herein, means oriented in parallel with the direction of thebiopolymer fibrils. In preferred embodiments, the living cells establishintercellular and extracellular connections such as those found in vivo.These connections can include, for example, intercalated disks betweendistal ends of adjacent cardiac myocytes consisting of gap junctionsthat mediate electrical signaling between cells, adherens junctions anddesmosomes between adjacent cells (cadherin interactions), and focaladhesions and hemidesmosomes (integrin-matrix interactions).

The living cells in the present invention are selected from the groupconsisting of myocyte precursor cells, cardiac myocytes, skeletalmyocytes, satellite cells, fibroblasts, cardiac fibroblasts,chondrocytes, osteoblasts, endothelial cells, epithelial cells,embryonic stem cells, hematopoetic stem cells, neuronal cells,mesenchymal stem cells, anchorage-dependent cell precursors, andcombinations thereof. For example, the living cells can be a combinationof cardiac myocytes and cardiac fibroblasts. In some embodiments theliving cells are contractile. As used herein the term “contractile”means having the ability to shorten in length due to mechanicalalterations in intracellular structural proteins. In preferredembodiments, the contractile cells are cardiac myocytes.

Neonatal cardiac myocytes maintain some plasticity, as they undergohyperplastic growth to reach maturity following birth. On a planargrowth surface, neonatal myocytes exhibit a stellate morphology unlikethat seen in vivo. Given more appropriate growth substrates in vitro,such as an oriented matrix of collagen or laminin, neonatal cardiacmyocytes will align with and attach to the matrix, adopt a rod-shapedmorphology, and form organized arrays of myofibrils. See i.e. McDevitt,T. C., et al., J Biomed Mater Res, 60:472-9 (2002). These cells can alsoform intercalated disks to allow for the conduction of electricalsignals from one cell to the next, resulting in coordinated contractileactivity similar to that seen in the intact myocardium of the heart.However, conventional 2-dimensional culture systems cannot sustain thisactivity indefinitely, and cells stop beating after two to three weeks.

The 3-dimensional structure of the scaffolding used in the presentinvention provides a more physiologically similar context than2-dimensional conditions, allowing, for example, grafted cardiacmyocytes to form intercellular and extracellular connections betweensuccessive layers, so that the tube can contract as an organ. Withoutbeing bound by this or any other theory, the inventors believe that theability of the ventricle of the heart to contract in a manner whichpropels blood out of the intraventricular space and into the systemiccirculatory system may rely on the unique geometric configuration of thecollagen fibrils in the extracellular matrix to properly align the cellsof the myocardium. Therefore, the purpose of the particular alignment ofbiopolymer fibrils of the present invention is to allow for thegeneration of the same type of mechanical force, given the introductionof contractile cells to the scaffold. Accordingly, in one embodiment ofthe present invention, the cardiac myocytes contract synchronously alongthe helical pattern of the biopolymer, thereby twisting the tube whilereducing its diameter, and thereby “wringing” fluid from the tube in amotion similar to that of a biological heart.

The term “contract synchronously” as used herein in reference to cardiacmyocytes, refers to the ability of an electrical signal to pass rapidlyfrom cell to cell via gap junctions to couple the cells so that theycontract in unison as a single functional unit.

In one embodiment, artificial tissue of the present invention has livingcells attached to two or more tubular tissue scaffolds, which maypreferably be in the form of concentric tubes. In some embodiments, theliving cells attached to individual tubular tissue scaffolds are ofdifferent cell types. By way of example, in artificial tissue comprisingthree concentric tubular tissue structures, the inner tube is seededwith contractile muscle cells and endothelial cells, the central tube isseeded with extracellular matrix-producing fibroblasts, and the outertube is seeded with contractile muscle cells and endothelial cells.

In another embodiment, the artificial tissue can be engineered tocontain and release cytokines and other pharmacologic agents that affectcell proliferation, development, migration, differentiation, and/oractivity, which are appropriate for the specific application of thetissue. As used herein, such cytokines and other pharmacologic agentsare referred to as “growth factors”, and include, without limitation,epidermal growth factor (EGF), fibroblast growth factor (FGF),erythropoietin (EPO), hematopoietic cell growth factor (HCGF),platelet-derived growth factor (PDGF), stem cell factors, bonemorphogenic protein (BMP), fibronectin, transforming growth factor-beta(TGF-β), and neurotrophins.

Growth factors modulate and control the inception, rate, and cessationof the vital healing events that may be associated with surgicalimplantation of engineered tissue. Growth factors are polypeptides thatmodulate cellular function and regulate cellular growth. These peptidesare extremely potent and, in very small quantities, are able to induce aspecific cellular response. Three key growth factors known to be vitalto the proper healing of damaged tissue are vascular endothelial growthfactor, platelet derived growth factor and nerve growth factor.

Vascular endothelial growth factor, VEGF, is an extracellular signalprotein that acts through the membrane bound tyrosine kinase receptor,VEGF receptor, to stimulate angiogenesis in vivo. A shortage of oxygenin practically any type of cell causes an increase in the intracellularconcentration of the active form of a gene regulatory protein calledhypoxia inducible factor I (HIF-1). HIF-1 stimulates transcription ofthe VEGF gene (and others). VEGF is secreted, diffuses through thetissue and acts on nearby endothelial cells. The response of theendothelial cells to VEGF includes at least four components. First, thecells produce proteases to digest their way through the basal lamina ofthe parent vessel. Second, the endothelial cells migrate toward thesource of the signal. Third, the cells proliferate. Fourth, the cellsform tubes and differentiate. Thus, endothelial cells create and linethe lumen of the newly formed blood vessels in the tissue. Theseneovessels will not persist on their own. They will developmicroaneurysms as well as other abnormalities that eventually rupture.These vessels rely on the recruitment of pericytes, from thevasculature, under the influence of signals from the endothelial cells,to further mature into competent blood vessels with the addition ofvascular smooth muscle cells and extracellular matrix. The recruitmentof pericytes, in particular, depends on PDGF secreted by endothelialcells.

Almost every cell in almost every tissue is located within 50-100 μm ofa capillary. In the case of wound healing, for example, there is a burstof capillary growth in the neighborhood of the damage to satisfy thehigh metabolic requirements of the repair process. Local infections andirritations also cause a proliferation of new capillaries most of whichregress and disappear when the inflammation subsides. In all of thesecases, the invading endothelial cells respond to signals produced by thetissue that they invade. The signals are complex, but a key part isplayed by VEGF.

Platelet derived growth factor (PDGF) is an extracellular signal proteinthat acts through the membrane bound tyrosine kinase receptors, PDGFreceptors α and β, to stimulate the survival, growth and proliferationof various cell types in vivo. PDGF stimulates chemotaxis andproliferation of fibroblasts and smooth muscle cells as well as collagensynthesis and collagenase activity.

Nerve growth factor (NGF) is an extracellular signal protein that actsthrough a membrane bound tyrosine kinase receptor, Trk A, to stimulatesurvival and growth of neurons. Cell growth and division can becontrolled by separate extracellular signal proteins in some cell types.Such independent control may be particularly important during embryonicdevelopment when dramatic changes in the size of certain cell types canoccur. Even in adult animals, however, growth factors can stimulate cellgrowth without affecting cell division. The size of a sympatheticneuron, for example which has permanently withdrawn from the cell cycle,depends on the amount of nerve growth factor secreted by the target cellit innervates. The greater the amount of NGF the neuron has access to,the larger it becomes. Concentrations of 250 ng/ml of NGF have beenshown to cause migration of neurons through collagen gels in vivo.

Therefore, when vascularization and innervation of the artificial tissueof the present invention is required, the presence of growth factorssuch as those described above is necessary for these processes toproceed in vivo.

In order to provide a useful amount of one or more growth factors, andto control their release from the artificial tissue of the presentinvention, the tissue scaffolds of the present invention can beincubated with a growth factor, or combination of growth factors, andallowed to absorb the growth factor(s). Alternatively, growth factorsmay be incorporated into the biopolymer solution prior to formation ofthe tubular tissue scaffold. It should be appreciated however, that thepresent invention is not limited by any particular method of treatingthe tissue scaffold with a growth factor, and the invention isapplicable to any such method now known or subsequently discovered ordeveloped. Growth factors useful in the present invention include, butare not limited to, vascular endothelial growth factor (VEGF), epidermalgrowth factor (EGF), fibroblast growth factor (FGF), erythropoietin(EPO), hematopoietic cell growth factor (HCGF), platelet-derived growthfactor (PDGF), nerve growth factor (NGF), transforming growth factors αand β (TGF-α and TGF-β), or combinations thereof.

In one embodiment, the tissue scaffold can be UV irradiated in order tocrosslink the biopolymer fibrils. Such irradiation can be administeredbefore or after the addition of growth factors. In one embodiment, thetissue scaffold is irradiated after the addition of growth factors, butprior to cell seeding in order to crosslink the biopolymer fibrils aftergrowth factor addition to the scaffold. While not wishing to be bound bythis or any theory, the degree of crosslinking can control the rate atwhich growth factors are released from the tissue. For example, thetissue scaffolds can be exposed to UV radiation at a wavelength of 254nm and an energy density of 500 μw/cm² for 140 minutes. When compared totissue scaffolds not subjected to UV radiation, tissue scaffolds treatedas described above have a slower rate of release of growth factors.Therefore, treatment of the tissue scaffolds with UV radiation can beused to modulate the release of growth factors once cells have beenseeded on the scaffold, in turn regulating processes such asneovascularization and innervation of the engineered tissue.

The present invention also encompasses a method of treatment comprisingimplanting in the body of a subject artificial tissue comprising livingcells attached to a tubular tissue scaffold comprising a tube having awall, wherein the wall includes biopolymer fibrils that are aligned in ahelical pattern around the longitudinal axis of the tube where the pitchof the helical pattern changes with the radial position in the tubewall.

The term “subject” for purposes of treatment includes any vertebrate.Preferably, the vertebrate is a human or animal subject who is in needof treatment for an injury, disease, or disorder of the type that can betreated by the use of artificial tissue. The subject is typically amammal. “Mammal”, as that term is used herein, refers to any animalclassified as a mammal, including humans, domestic and farm animals, andzoo, sports, or pet animals, such as dogs, horses, cats, cattle, etc.Preferably, the mammal is a human. Additionally, the term “implanting inthe body” refers to surgically inserting into the subject at the site ofthe injury, disease, or disorder being treated.

When it is said that the present artificial tissue can be used intreatment for an injury, disease, or disorder of the type that can betreated by the use of artificial tissue, such treatment can include, butis not limited to, replacement of vessels such as, for example, coronaryarteries; repair and/or replacement of any other physiologic tubularstructures, such as, for example, ureters, veins, lymph channels, GItract components, and the like; repair of injured bone; repair ofdamaged nervous tissue as in, for example, spinal cord injury; orcorrection of impaired cardiac function caused by, for example, ischemicheart disease.

In one embodiment, the artificial tissue used in the method of treatmenthas living cells introduced to the luminal and outer surfaces of thetissue scaffold, and preferably the cells are aligned along the helicalpattern of the biopolymer. The living cells in the present invention canbe selected from the group consisting of myocyte precursor cells,cardiac myocytes, skeletal myocytes, satellite cells, fibroblasts,cardiac fibroblasts, chondrocytes, osteoblasts, endothelial cells,epithelial cells, embryonic stem cells, hematopoetic stem cells,neuronal cells, mesenchymal stem cells, anchorage-dependent cellprecursors, and combinations thereof. In some embodiments, the livingcells are selected from the group consisting of cardiac myocytes,skeletal myocytes, fibroblasts, cardiac fibroblasts, chondrocytes,osteoblasts, endothelial cells, epithelial cells, embryonic stem cells,hematopoetic stem cells, neuronal cells, mesenchymal stem cells andcombinations thereof.

In one embodiment, the living cells of the present invention originatefrom the subject receiving treatment. Autologous grafts are far lesssusceptible to rejection, as they are not recognized as foreign and,hence, do not elicit an immune response in the subject. In preferredembodiments, cells taken from the subject are introduced to the tissuescaffold in vitro. The cells are cultured for the period of timenecessary for the cells to degrade the original scaffold biopolymer andregenerate a matrix of secreted extracellular matrix proteins.Preferably, the cells are cultured in vitro for a period of timenecessary for the cells to degrade and replace at least 25 percent ofthe original scaffold. More preferably, the cells are cultured for aperiod of time necessary for the cells to degrade and replace at least50 percent of the original scaffold and, more preferably still, a periodof time necessary for the cells to degrade and replace at least 75percent of the original scaffold. The remodeled tissue scaffold is thenconstructed of proteins produced by the subject's own cells, and istherefore not recognized as a foreign substance that would induce animmune response in the subject.

It is preferred, in the present invention, that the living cellsintroduced to the tubular tissue scaffold establish intercellular andextracellular connections such as those found in vivo. In oneembodiment, these living cells are contractile. In preferred embodimentsthe contractile cells are cardiac myocytes. As mentioned previously, theunique organization of the tissue scaffold of the present invention isable to guide the formation of an interconnected, contractile cellsystem, wherein the cardiac myocytes contract synchronously along thehelical pattern of the biopolymer in a “wringing” motion to pump fluidthrough the lumen of the tubular tissue scaffold. The pumping action ofthe artificial tissue can be made to be directional by any one ofseveral methods. For example, one-way valves can be placed on one orboth sides of a contractile artificial tissue tube that result indirecting the flow of fluid through the tube. Alternatively, one or moreartificial tissue tubes can be located in series and controlled with oneor more pacing devices in a manner that creates a peristaltic action toforce fluid from one end of the tubular construct to the other.

Because the artificial tissue is able to act as a pump when contractilecells are present, this particular embodiment of the present inventioncan be used in the treatment of an injury, disease, or disorder thatinvolves abnormalities in either cardiac output or vascular tone, orvascular blockage. In some embodiments of the present invention, thesecan include congestive heart failure, dilated cardiomyopathy,hypertrophic cardiomyopathy, infiltrative cardiomyopathy, ischemic heartdisease, heart attack, heart failure, coronary artery disease,atherosclerosis, hypertension, chronic renal disease, cerebrovasculardisease, carotid artery disease, and peripheral vascular disease.

For most in vitro or in vivo uses, the tubular scaffold used to make theartificial tissue of the present invention should be sterilized andtreated with antibiotic and antifungal agents. By way of example, thetubular scaffolds can be sterilized by placing them in sterile Mosconassolution (0.14M NaCl, 0.0027M KCl, 0.012M NaHCO₃, 4.2×10⁻⁵M NaH₂PO₄,0.0094M glucose) and exposing them to gamma radiation or ultraviolet(UV) radiation or both for up to 4 hours. Following UV treatment, freshMarconas solution with 0.01 mg/ml gentamycin, 4 μg/ml Amphotericin-B,and 10 μg/ml fibronectin can be added to a culture dish containing thetubular scaffolds and the tubes can be incubated for 24 hours with 5%CO₂ at 370 Celsius.

The artificial tissue of the present invention provides a tubularcell-based prosthetic that would be particularly useful for repairing orreplacing tube-shaped “organs”, such as various sized blood vesselsincluding coronary arteries and renal arteries, ureters, fallopiantubes, or nerve fiber conduits. If the cells of the present inventioninclude contractile cells, such as cardiac myocytes, the tube as a wholecan contract in a “wringing” motion, as dictated by the alignment of thecells along the helical pattern of the biopolymer fibrils. This type ofcontractile engineered construct could be used, for example, as aventricular assist device to enhance cardiac contractility in a failingheart. Large versions of such a contractile tubular artificial tissuecould be implanted near the heart to improve systemic circulation.Smaller versions could be implanted near vital organs to improve localperfusion. For example, contractile tubes could be transplanted in therenal arteries to increase renal perfusion and improve cardiacperformance while unloading the heart. In another example, contractiletubes could be implanted in the common carotid artery inferior to thecarotid sinus to improve brain perfusion. The contractile tubes could beused as replacement coronary arteries to provide both a blood conduit aswell as improved local perfusion during diastole. Preferably, thecontractile activity of this type of construct is controlled with apacing device. As used herein, “pacing device” refers to any electricaldevice that can be used to maintain a particular rate of contraction,such as, for example, the implantable cardiac resynchronization devicemade by Medtronic, Inc., called InSync®.

One problem that has limited the application of artificial tissue, andin particular, artificial muscle tissue such as replacement cardiacmuscle tissue, is that the tissue must be able to function immediatelyand accurately in the rigorous environment of the continuouslycyclically-contracting heart. Such tissue cannot be allowed tosignificantly degrade and be remodeled in situ. Currently availablescaffolds for tissue engineering undergo global remodeling onceimplanted in vivo. In order to avoid or reduce the degree of remodelingand consequent loss of strength and/or function of implanted artificialtissue, the present invention includes a method for preconditioning thenovel tubular artificial tissue scaffold by seeding it with cells, suchas fibroblasts, and stimulating the cells with a combination of matrixcomponents, growth factors, and mechanical stimulation. Without beingbound by this theory, it is believed that such stimulation can controlfibroblast proliferation, matrix synthesis, matrix degradation andfibroblast apoptosis, among other parameters that are important forartificial tissue development.

When it is said that the preconditioning includes stimulation of thecells with matrix components, what is meant is the selection and controlof the composition of the present artificial tissue scaffold and itsphysical structure, as have been discussed above. For example, variablessuch as the type of biopolymer that is used to produce the fibrils, andhow and to what extent they are aligned, as well as the dimensions ofthe tubular scaffold, are included in the selection and control ofmatrix components.

Mechanical signals play an integral role in both directing myocytes toassume the distinctive cytoarchitectural features characteristic ofdifferentiated myocytes as well as the arrangement of individualmyocytes into an intact muscle. A primary developmental force in theheart is mechanical signaling resulting from contractile force as wellas an increase in pressure and volume. See Terracio, L., et al., InVitro Cell Dev Biol, 24:53-8 (1988). The regulation, expression,synthesis, and degradation of various contractile and regulatoryproteins, as well as cell size, are influenced by mechanical stress.Accordingly, in one embodiment of the present invention, artificialtissue containing certain cells, such as contractile cardiac myocytes,is preconditioned by the application of mechanical stress before beingimplanted in the subject. What is meant by “preconditioned by theapplication of mechanical stress” is permitting cells to becomeaccustomed to forces, such as stretch, normally found in thecardiovascular system in vivo, by subjecting the cells to these forcesbefore introduction into the subject.

One method of applying mechanical stimulation to the artificial tissuescaffold is by culturing artificial cardiac myocyte-containing tissue ina pulsatile internal flow stretch bioreactor that mimics the action of adeveloping cardiovascular system. Stretch has been shown to up-regulateboth matrix metalloproteinase production as well as the synthesis of newmatrix. It is believed that the pulsatile internal flow and stretchmimics the action of a developing cardiovascular system.

An example of a pulsatile internal flow stretch bioreactor is shown inFIG. 8. In this system, a tubular tissue scaffold, either before orafter the attachment of cells is placed in the flow pattern of thebioreactor. The media reservoir can be adjusted to various heights fromessentially zero height to a height equivalent to 10 cm of waterpressure, or more. The pulse frequency and stretch frequency can beadjusted from 0 Hz to 2 Hz. It is believed that fibroblasts and myocytesthat are seeded on the tubular tissue scaffold will remodel the matrixin response to the strain imposed by the pulse.

In addition to the presence of growth factors during the preconditioningprocess, other biologically active chemicals, such as matrixmetalloproteinase (MMP) inhibitors can also be added in order to promotethe preconditioning of the artificial tissue.

In one embodiment, the method for preconditioning the novel tubularartificial tissue scaffold comprises seeding a tubular artificial tissuescaffold having aligned biofibrils with living cells and culturing thecells in the presence of media containing at least one growth factor andunder conditions where the tubular artificial tissue scaffold issubjected to stretch and pressure pulse of controlled frequency andamplitude.

The present invention also includes a novel preconditioned artificialtissue comprising living cells that are attached to a tubular artificialtissue scaffold having aligned biofibrils, wherein the cells have beencultured in the presence of media containing at least one growth factorand under conditions where the tubular artificial tissue scaffold issubjected to stretch and pressure pulse of controlled frequency andamplitude.

In another embodiment of the present invention, the tubular artificialtissue is split along the longitudinal axis and opened to form a sheet.The sheet contains layers of aligned biopolymer fibrils with attachedcells, where the direction of the alignment changes in each successivelayer. Applications for this type of artificial tissue could include,for example, artificial skin to treat burn injury or surgical patchesfor internal application. Cardiac myocytes seeded on a collagen sheethaving the fibril alignment patterns of the present novel tissuescaffolds could provide functional myocardial patches to treat infarctedareas of the heart.

In one embodiment, the tubular tissue scaffold is split longitudinallyand opened into a sheet prior to attaching living cells to the tissuescaffold. The sheet can be coated with an additional layer of biopolymerfibrils such as, for example, collagen, wherein the biopolymer filbrilspolymerize and are oriented in the direction in which they are applied.

Living cells useful in this embodiment of the present invention include,for example, myocyte precursor cells, cardiac myocytes, skeletalmyocytes, satellite cells, fibroblasts, cardiac fibroblasts,chondrocytes, osteoblasts, endothelial cells, epithelial cells,embryonic stem cells, hematopoetic stem cells, neuronal cells,mesenchymal stem cells, anchorage-dependent cell precursors, andcombinations thereof. In preferred embodiments, the living cells areselected from the group consisting of myocyte precursor cells, cardiacmyocoytes, skeletal myocytes, satellite cells, fibroblasts, andcombinations thereof. In one embodiment, the living cells originate fromthe subject receiving treatment.

The sheet form of the tissue scaffold can be used for treatment of, forexample, hernia, heart attack, congenital heart defects, tissue missingdue to congenital defect, skin burns, organ damage, and muscle damage.

In one embodiment, the sheet form of artificial tissue can be used torepair a tear or defect in a variety of tissues including, but notlimited to, cardiac tissue, skeletal muscle tissue, epithelial tissue,vascular tissue, nerve tissue, lymphatic tissue, connective tissue,epidermal tissue, endocrine tissue, cartilage, and bone.

By way of example, the sheet form of artificial tissue can be used torepair a ventral hernia. Collagen tubes are produced as describedherein, and are cut open longitudinally to form sheets. A thin layer ofcollagen solution can be streaked on the surface of the substrates andallowed to polymerize. This procedure results in a thin layer ofcollagen fibrils that are arrayed in parallel with one another along thedirection that the collagen solution was streaked. Myocyte precursorcells such as, for example, satellite cells, can be introduced to thesheet scaffold and cultured until the cells diffentiate into myocytesand fuse to form mature myotubes. This artificial skeletal muscle tissuecan then be surgically applied to the damaged area of the abdominal wallto repair the hernia.

Additionally, the present invention embraces a method of identifying theeffects of a pharmaceutical composition on cell function comprisingadministering said pharmaceutical composition in vitro to artificialtissue comprising living cells attached to a tubular tissue scaffoldcomprising a tube having a wall, wherein the wall includes biopolymerfibrils that are aligned in a helical pattern around the longitudinalaxis of the tube where the pitch of the helical pattern changes with theradial position in the tube wall. Artificial tissue that has been madeinto sheet form as described above can also be used for the tissuescaffold for this method.

In one embodiment, the method further comprises determining the effectsof a pharmaceutical composition on the cells by measuring or identifyingchanges in cell function. This can be accomplished by many methodologiesknown to those skilled in the art including, for example, Western blotanalysis, Northern blot analysis, RT-PCR, immunocytochemical analysis,BrdU labeling, TUNEL assay, and assays of enzymatic activity. In someembodiments, the living cells are contractile cardiac myocytes.Accordingly, measurements of parameters such as isovolumic pressuregeneration, length tension, and isometric force generation can be made.By way of example, the instant artificial tissue containing cardiacmyocytes that contract in culture could be treated with an agent, suchas ephedrine, and any change in contractility of the myocytes can bemeasured as described in Example 3. This method provides an in vitrodiagnostic system that can be utilized to rapidly assay thephysiological consequences of administration of a given pharmaceuticalcomposition on cell function such as, for example, cardiaccontractility.

The following examples describe preferred embodiments of the invention.Other embodiments within the scope of the claims herein will be apparentto one skilled in the art from consideration of the specification orpractice of the invention as disclosed herein. It is intended that thespecification, together with the examples, be considered to be exemplaryonly, with the scope and spirit of the invention being indicated by theclaims which follow the examples. In the examples all percentages aregiven on a weight basis unless otherwise indicated.

EXAMPLE 1

This example illustrates the preparation of a biopolymer gel comprisingtype I collagen from a bovine steer hide.

Collagen was prepared from the hide of an 18-month old bovine steer byremoval of the superficial epidermis including the hair and folliclepits. To isolate the collagen the hide was cut into 4×6 cm strips andwashed with three changes of deionized (DI) water for 1 hr. each (thesecond wash contained 0.2% NaHCO₃). The remaining follicles andnoncollagenous proteins were removed by bathing in a solution of 0.6%NaHCO₃, 2% Ca(OH)₂, and 4.3% NaHS for 30 min at 20° C. After threewashings in DI water, the strips were treated overnight in 2% Ca(OH)₂ at4° C. The remaining fat and epidermis were trimmed from the strips thefollowing day. The strips were then placed in a 2M NaCl solution andneutralized with HCl to a pH of 6.8-7.0. The strips were then washedthree times in DI water, cut into 0.5×2 cm pieces, and placed in asolution of 0.5 N acetic acid with or without pepsin, 1:100 based onhide weight. After incubating overnight at 4° C., the swollen stripswere mixed with ice and emulsified into a gel dispersion using aKitchen-Aid food processor (model # FP500WH; St. Joseph, Md.). Thesuspension was centrifuged at 9950×g for 35 min to remove small,unsolubilized particles. Type I collagen was precipitated from the geldispersion by mixing with NaCl to a final concentration of 2M. Thecollagen was collected by centrifugation at 9950×g for 35 min,resuspended in DI water and neutralized to a pH of 7.2 using NaOH. Thecollagen suspension was then dialyzed vs phosphate buffered salineovernight and vs DI water for three changes over 24 hr. The resultingcollagen gel was then centrifuged at 9950×g for 35 min to remove excesswater and the collagen concentration adjusted to 25 mg/ml by addition ofDI water and the pH adjusted to 2.5-3.0 with concentrated HCl.

Example 2

This example illustrates the production of a collagen tubular tissuescaffold having aligned fibrils by using a counter-rotating coneextruder.

Collagen tubes were produced by loading a collagen gel dispersion,produced as described in Example 1, into a 60 ml syringe that was placedinto a syringe pump (having an adjustable feed rate of 0-20 cc/min) thatfed the collagen into the feed port of a counter-rotating cone extruderthat was fabricated from Delrin (Tri Star plastics, Reading, Pa.)according to the illustration shown in FIG. 1. The drive motor was a 90volt 0.06 hp DC motor (Baldor Electric, Ft. Smith Ak.) with aregenerative feedback drive and a 5 kΩ potentiometer for user interface(KB electronics, Coral Springs Fla.). The motor is coupled to the piniongear which drives the outer rotating member and the inner rotating conedrive gears. The direct drive system can control speeds from 0 to 1800rpm with a 2:1 reduction from the pinion gear to the drive gears. Theextruder was mounted atop a controlled atmosphere chamber that consistedof a vented acrylic box that was purged at 1 L/min with a 50/50 mix ofanhydrous ammonia and air.

As the dispersion was fed between the outer and inner rotating cones, acollagen tube exited through the annular-shaped exit port and wascollected in a deionized water bath. The speed of the cone rotation andforward flow of the gel dispersion combined to create a helicallyaligned orientation pattern of the collagen fibrils. To maintain alumen, as tubes are extruded, gas (air or air and ammonia) is meteredbetween 10 to 60 ml/min through the tube. Following extrusion the tubewas left in a water bath for 1 hr at room temperature followed by 30 minin a 0.3% NaHCO₃ solution and an additional 1 hr. in deionized water.

Determination of the Pitch of the Helical Pattern of the AlignedCollagen Fibrils: By Polarized Light Microscopy:

To determine the pitch of the collagen fibrils at the outside andluminal walls of the tube, freshly extruded collagen tubes were fixed in2.5% glutaraldehyde in phosphate buffered saline (PBS) for 1 hr at 25°C. The tubes were then stained with picrosirius red, and then imagedwith a Bausch & Lomb Illuminator optical polarized microscope to examinethe birefringent nature of the collagen fibrils. The lower polarizer wasinstalled such that light was polarized in a north-south direction andthe upper analyzer was oriented in the east-west direction.

Samples were placed on the rotary stage of the polarized lightmicroscope in an east-west direction with respect to the long axis ofthe tube. The stage was rotated from 0° to 180° and the stage angleshowing maximum darkness and brightness recorded. When the collagenfibers were at a 45° angle between the polarizer and the analyzer,transmitted light was at its maximum intensity and the collagenorientation could be directly visualized. When the pitch of the helicalpattern coincided with either the lower polarizer or upper analyzer, therefracted light was quenched and there was minimal light transmission.The rotation angle of the stage was recorded at each interval of maximumand minimum light transmission and the angle of the fiber arraycalculated with respect to the long axis of the tube.

By Scanning Electron Microscopy (SEM):

To further illustrate the helical orientation of the collagen fibrils inthe tubes, samples were prepared for SEM by the O-GTA-O-GTA-O method ofTakahashi et al., J. Electron Microsc., 35(3):304 (1986). Extrudedcollagen tubes were fixed for 2 hr. in 3% glutaraldehyde in 0.1M sodiumcacodylate buffer (pH 7.4), rinsed in buffer, and immersed in 2% aqueousOsO₄ for 2 hr. After rinsing, samples were treated with two repetitionsof GTA-0 steps: 3×4 hr 8% glutaraldehyde/2% tannic acid at 4° C., rinse,2 hr. 2% OsO₄. Tubes were then dehydrated in a graded ethanol, criticalpoint dried, mounted on aluminum stubs and imaged on a JEOL JSM-6300V at10 kV.

The counter rotating action of the extruder cones created a spiraling,or helical, alignment of the collagen fibrils that had a uniformdirection on each of the outside and luminal walls of the tube. Asection of a tube was split longitudinally down the tube wall and thetube was opened into a sheet. SEM microscopy of each side of the sheetis shown in FIG. 2 and clearly shows the “corkscrew” and “countercorkscrew” orientation of the fibrils at the surface of the outside andluminal tube walls. In this tube the pitch of each helix was about 45°,and in opposite directions.

When the cone rotation rate that varied from 150 rpm to 900 rpm at afixed extrusion rate of 150 cm/min, the fibril angle of each array withrespect to the long axis of the tube varied between 26° to 62°. Slowerrotation speeds resulted in smaller angles while higher cone rotationspeeds resulted in larger angles.

Example 3

This example illustrates seeding of cardiac myocyte on a collagentubular tissue scaffold.

Collagen tubes were produced as described in Example 2. Prior to theaddition of myocytes, the collagen tubes were sectioned into 1.25 cmlengths. The small collagen tube sections were placed in sterileMosconas solution (0.14M NaCl, 0.0027M KCl, 0.012M NaHCO₃, 4.2×10⁻⁵ MNaH₂PO₄, 0.0094M glucose) and exposed to ultraviolet (UV) light for 4hours to sterilize. Following UV treatment, fresh Marconas solution with0.01 mg/ml gentamycin, 4 μg/ml Amphotericin-B, and 10 μg/ml fibronectinwas added to a culture dish containing the collagen tube sections andincubated for 24 hours with 5% CO₂ at 37° Celsius.

Isolation of neonatal cardiac myocytes was performed as described bySimpson et al., J Cell Physiol, 161:89-105 (1994). For seeding, collagentube sections were placed in 150 mm culture dishes and 0.5 ml of acellular suspension containing 2×10⁶ cells/ml was injected into thelumen of each tube using an 18 gauge IV catheter (Surflo Terumo Mol.Corp., Somerset, N.J.). The tubes were then placed in a SyntheconRotating Wall Bioreactor (Synthecon, Houston, Tex.) and the reactorfilled with an additional 2.5 ml of cell suspension. Incubation was at arotation rate of 20 rpm with 5% CO₂ at 370 Celsius for 72 hours.

Following incubation in the bioreactor, tubes were placed in culturedishes with fresh media containing 4 μg/ml cytosinep-D-arabino-furanoside. After 72 hours in the bioreactor, individualmyocytes contracted spontaneously. Two further cell seedings wereperformed by adding 2×10⁶ cells per tube to the tube lumen and outsidesurface at one week intervals. After 10 days in culture, entire areas ofthe tube contracted and, after 16 days in culture, synchronous myocytecontraction and forceful contractions of the entire tube were observed.

Characterization of Cardiac Myocytes in Collagen Tubes:

For confocal microscopy, tubes were sectioned longitudinally, fixed andstained for f-actin and connexin 43 (Chemicon MAB3068; 1:1000 dilution)as described by Price, R. L., et al., Anat Rec, 245:83-93 (1996); andAngst, B. D., et al., Circ Res, 80:88-94 (1997). FIG. 5 demonstrates aconnexin 43 protein expression pattern similar to that described invivo, where connexin 43 is distributed along the periphery of thecardiac myocyte and in cell-cell junctions (A). For stereo imaging,Z-series were collected at 1 μm intervals to a maximum depth of 80 μm.Reconstructed confocal microscopy Z-series through cardiac myocyteseeded tubes after 4 weeks in culture, shown in FIGS. 5 (B) and (C),show four to five layers of cardiac myocytes aligned in parallel withthe collagen fibrils in the tube wall.

Cardiac myocyte seeded tubes were processed for transmission electronmicroscopy (TEM) as previously described in Price, R. L., et al., AnatRec, 245:83-93 (1996). TEM images, like those in FIGS. 6 (A) and (B),demonstrate the presence of organized, aligned myofibrils (F),well-developed sarcomeres (not shown), Z-bands (Z), numerousmitochondria (M), cell:cell junctions (arrows in inset A), andattachment to collagen fibrils (arrows in inset B), characteristic ofthe cardiac myocyte phenotype seen in vivo.

Determination of Electrical Activity of Cardiac Myocyte Seeded Tubes:

Seeded tubes were measured for electrical activity by attaching twoplatinum wires at opposite ends of the tube, using a third wire as areference in the surrounding media. The wires were shielded frominterference by enclosing all but the distal end of each wire within aplastic casing. These electrodes were connected to an amplifier, ananalog-to-digital converter, and finally to a computer where the signalwas recorded. One-hundred seconds of electrical activity was collectedfor eight different tubes at 100 samples per second. The signal was thenprocessed using a Fast Fourier transform algorithm to show theelectrical activity of individual tubes. FIG. 7 (A) is a representativeelectrical signal recorded from a tube after 4 weeks in culture, andFIG. 7 (B) illustrates the Fast Fourier transform analysis of theelectrical data, demonstrating a dominant frequency at 3.4 Hz. Thiscorresponds to the contraction frequency of about 200 beats per minuteobtained by visual observation and counting of contractions.

Example 4

This example demonstrates the use of the disclosed tissue scaffold forventral hernia repair.

Collagen tubes can be made following the procedure described in Example2, or according to any of the methods set forth in this disclosure. Thecollagen tubes are cut open longitudinally to form sheets, and a thinlayer of collagen solution is streaked on the surface of the substratesand allowed to polymerize. This procedure results in a thin layer ofcollagen fibrils that are arrayed in parallel with one another along thedirection that the collagen solution was streaked. Skeletal myoblastswere plated on the patterned collagen sheets. Over several days, themyoblasts continued to fuse and a series of uniformly arrayed, denselypacked myotubes with morphology reflective of skeletal muscle developed.

The majority of the cells placed on the aligned collagen are MyoDpositive myoblasts that fuse into multinucleate myotubes. Confocalmicroscopy of a phalloidin stained culture demonstrates parallel-alignedskeletal muscle cells with uniformly spaced sarcomeres. Located betweenthe skeletal muscle cells are a population of fibroblast-like cells thatare a mixture of myoblasts and fibroblasts that can produce collagen. Toincrease the thickness of the aligned skeletal muscle cultures and forma more tissue-like construct, the collagen sheets with aligned myotubesare placed in the Rotating Wall Bioreactor (RCCS) (Synthecon), andadditional cells are seeded to acheive multilayering. It is possible toobtain as many as 40 layers of skeletal muscle cells in the RCCS.

In order to use these constructs for repair of ventral hernia, anexperimental model of hernia in a rat is used. After achievement ofgeneral anesthesia, the animal's abdomen is opened and a 1 cm² midlineabdominal wall defect is created. At this point, the engineered skeletalmuscle tissue is surgically sewn into the defect and the abdomen isclosed. Topical antimicrobials are placed on the incision and the animalis allowed to ascend from anesthesia. Intramuscular analgesics areadministered as well as subcutaneous fluid boluses while the animalrecovers. Tissue integration is monitored post operatively by clinicallyinspecting the surgical sight for signs of infection, bleeding, suturefailure, and extrusion of viscera through the rectus abdominous,indicative of repair failure. Animals are sacrificed at specific timepoints and the tissue is evaluated for specific events known to occurthroughout the wound healing process.

The implanted tissue is evaluated microscopically, monitoring normalwound healing processes, neovascularization, innervation, inflamationand infection. Inflammatory cells, tissue morphology and detection ofangiogenesis are investigated using hematoxylin and eosin (H&E) stains.For this evaluation, 5 mm diameter punch biopsies of the interfacebetween the native tissue and the repair material are obtained andserially sectioned. The sections are evaluated for clinicallysignificant signs of rejection or infection such as extensiveinterstitial mononuclear cell infiltration and edema as well as mildinterstitial hemorrhage and antibody-mediated rejection by the presenceof significant foamy macrophage populations.

Neovascularization should be apparent on the H&E slides. Endothelialcells as well as the internal elastic lamina and vascular smooth musclecells are identified for arterial vessels. The tunica intima, tunicamedia and vasa vasorum are identified for venous vessels. The presenceof red blood cells within the vessels is a strong indication of theirfunctionality. Collagen accumulation and location are evaluated usingMasson's trichrome staining. Myoneural junctions should clearly displaythe myelinated nerve fiber approaching the skeletal muscle fiber. As theaxon nears the muscle cell, it loses its myelin sheath and continues onas a non myelinated axon which can be visualized by Sevier-Mungerstaining for neural tissues. Motor end-plates should also be apparentusing Gwyn-Heardman staining method described in the Color Atlas ofHistology, 3^(rd) ed. Baltimore: Lippincott Williams and Wilkins (2000).

The tissue at the repair site is also analyzed immunohistochemically forthe temporal and spatial localization of growth factors VEGF, PDGF, andNGF within and around the repair. Two sets of the 2 mm biopsies aretaken. One set is analyzed for growth factors by enzyme-linkedimmunosorbent assay (ELISA) to quantify the concentration of growthfactors within the tissue (ELISA kits: VEGF and PDGF from R& D systems,NGF from Chemicon). The second set is fixed with paraformaldehyde,sectioned, and immunohistochemically stained for VEGF, PDGF and NGFusing commercially available antibodies (R&D systems). The antibodystaining is imaged using a laser scanning confocal microscope tolocalize the concentrations of growth factors spatially within thetissue using the methods of Germani. A., et al., Am J. Path.163(4):1417-1428 (2003). Briefly, samples are fixed in 4%paraformaldehyde overnight at 4° C. They are vibratome sectioned andstained with 488-phalloidin for f-actin, and immunohistochemicallylabeled using commercially-available antibodies. Serial sections arelabeled with either VEGF, PDGF or NGF, one growth factor label persection. Secondary antibodies are labeled with Texas Red.Immunohistochemical visualization is performed using a Bio-Rad 1024 ESlaser scanning confocal microscope.

Additionally, the tissue is evaluated mechanically for normal woundhealing repair strength development. The indenter force required topuncture the abdominal repair is measured using an Instron® forcemeasurement system. Measurements are made on the abdominal repair byresecting the entire abdominal wall free from the animal and placing itin ice cold Krebs Ringer solution with 30 mM 2,3-butanedione monoxime(BDM) to prevent dissection injury. Samples are mounted in a bathperfused with warm (37° C.) Krebs Ringer solution oxygenated by bubblingthe solution reservoir with a 95% O₂/5% CO₂ mixture. The sample isplaced in the stainless steel “picture frame” sample holder which clampsthe sample along the edges in a plane parallel to the base of theInstron® and perpendicular to the indenter. A 3 mm diameter sphericalindenter is used to impinge on the center of the repair site. The sampleis indented at a rate of 3.0 mm/sec. The force and displacement of theindenter and the peak force required to burst through the repair siteare measured and recorded. Samples for mechanical testing are fromanimals that have not undergone punch biopsies. Data will be analyzedusing ANOVA followed by pair-wise comparison.

All references cited in this specification, including without limitationall papers, publications, patents, patent applications, presentations,texts, reports, manuscripts, brochures, books, internet postings,journal articles, periodicals, and the like, are hereby incorporated byreference into this specification in their entireties. The discussion ofthe references herein is intended merely to summarize the assertionsmade by their authors and no admission is made that any referenceconstitutes prior art. Applicants reserve the right to challenge theaccuracy and pertinency of the cited references.

In view of the above, it will be seen that the several advantages of theinvention are achieved and other advantageous results obtained.

As various changes could be made in the above methods and compositionswithout departing from the scope of the invention, it is intended thatall matter contained in the above description and shown in theaccompanying drawings shall be interpreted as illustrative and not in alimiting sense.

1. A method of producing a tubular tissue scaffold, the methodcomprising: providing a gel dispersion comprising a biopolymer; feedingthe gel dispersion to a tube-forming device that is capable of producinga tube from the gel dispersion while providing a radial shear forceacross the wall of the tube and having a gas channel connecting a gassource with the luminal space of the tubular tissue scaffold as it exitsthe tube-forming device; forming the gel dispersion into a tube; andcausing the gel dispersion to solidify, thereby forming a tubular tissuescaffold comprising a tube wall having biopolymer fibrils that arealigned in a helical pattern around the longitudinal axis of the tubeand where the pitch of the helical pattern changes with the radialposition in the tube wall.
 2. The method according to claim 1, whereinthe method comprises: providing a gel dispersion comprising abiopolymer; feeding the gel dispersion to a counter-rotating coneextruder having an annular-shaped exit port and having a gas channelconnecting a gas source with the center of the annular-shaped exit portand opening into the luminal space of the tubular tissue scaffold as itexits the extruder; extruding the gel dispersion to form a tube; andcausing the gel dispersion to solidify, thereby forming a tubular tissuescaffold comprising a biopolymer having fibrils in the tube wall thatare aligned in a helical pattern around the longitudinal axis of thetube and where the pitch of the helical pattern on the luminal surfaceof the tube is different from the pitch of the helical pattern on theoutside surface of the tube.
 3. The method according to claim 2, whereinthe biopolymer comprises collagen, fibronectin, laminin, elastin,fibrin, proteoglycans, hyaluronan, or combinations thereof.
 4. Themethod according to claim 3, wherein the biopolymer comprises collagen.5. The method according to claim 4, wherein the biopolymer comprisestype 1 collagen.
 6. The method according to claim 5, wherein the geldispersion of collagen is prepared by a method comprising: removing thehair from the hide of a bovine; washing the hide sequentially in water,water containing NaHCO₃ and surfactant, and water; contacting the hidewith an aqueous solution containing NaHCO₃, Ca(OH)₂, and NaHS; washingthe hide in water; treating the hide with an aqueous solution ofCa(OH)₂; rinsing the hide with water and trimming the hide of anyremaining skin tissue and fat; placing the hide in an aqueous NaClsolution and adding hydrochloric acid solution until the pH is stablebetween about 6.0 and 8.0; washing the hide in water; placing the hidein an aqueous solution of acetic acid with or without pepsin; mixing andallowing the hide to swell; placing the swollen hide in a mill andprocessing into a gel dispersion; filtering the gel dispersion to removeundissolved particles; centrifuging the gel dispersion to remove smallundissolved particles; adding NaCl to the gel dispersion in an amountsufficient to precipitate collagen from the gel dispersion; filteringthe collagen precipitate and resuspending it in deionized water; addingNaOH to bring the pH of the collagen dispersion to a pH between about 6and about 8; dialyzing the collagen dispersion against phosphatebuffered saline solution; resuspending the collagen in deionized water;centrifuging the collagen dispersion to concentrate solid collagen gelas a pellet; and resuspending the pellet in aqueous HCl.
 7. The methodaccording to claim 4, wherein the gel dispersion of collagen is preparedby a method comprising: removing the hair from the hide of a freshlykilled bovine; washing the hide sequentially in water, water containing0.2% NaHCO₃ and 0.02% surfactant, percent by weight, and water;contacting the hide with an aqueous solution containing 0.6% NaHCO₃, 2%Ca(OH)₂ and 4.3% NaHS, percent by weight, for a period of from about 0.1hrs to about 10 hrs; washing the hide in three changes of deionizedwater; treating the hide with an aqueous solution of 2% Ca(OH)₂, percentby weight, for a period of from about 1 to about 24 hours; rinsing thehide with water and trimming the hide of any remaining skin tissue andfat; placing the hide in an aqueous 2M NaCl solution and addinghydrochloric acid solution until the pH is stable between about 6.8 and7.0; washing the hide in three changes of deionized water; placing thehide in an aqueous solution of 0.5 N acetic acid with or without pepsinadded to give a 1:100 weight ration based on wet hide weight; mixing fora period of from about 1 to about 24 hours and allowing the hide toswell; placing the swollen hide in a high-shear mill with ice andprocessing into a gel dispersion; filtering the gel dispersion to removeundissolved particles; centrifuging the gel dispersion to remove smallundissolved particles; adding NaCl to the gel dispersion in an amountsufficient to give a 2 M concentration, which will cause theprecipitation of collagen from the gel dispersion; filtering thecollagen precipitate and resuspending it in deionized water; adding NaOHto bring the pH of the collagen dispersion to pH 7.2; dialyzing thecollagen dispersion against phosphate buffered saline solution for aperiod of from about 1 to about 24 hours; resuspending the collagen inthree changes of deionized water over a period of about 24 hours;centrifuging the collagen dispersion to concentrate solid collagen gelas a pellet; and resuspending the pellet in 0.012 N HCl to a finalcollagen concentration of about 20 mg/ml.
 8. The method according toclaim 2, further comprising feeding a gas from the gas source to theluminal space of the tubular tissue scaffold as it exits the extruder.9. The method according to claim 8, wherein the gas comprises a mixtureof air and ammonia gas.
 10. The method according to claim 9, wherein themixture of air and ammonia is about a 50:50 mixture by volume.
 11. Themethod according to claim 8, wherein the outside surface of the tubulartissue scaffold is contacted with a mixture of air and ammonia gas as itexits the extruder.
 12. The method according to claim 11, wherein thetubular tissue scaffold exiting the extruder falls into a liquid bathlocated to receive the tubular tissue scaffold from the extruder. 13.The method according to claim 12, wherein the liquid bath is located adefined distance beneath the exit port of the extruder.
 14. The methodaccording to claim 13, wherein the liquid bath is a aqueous ammonia bathat a pH of about
 10. 15. The method according to claim 13, wherein thebiopolymer gel dispersion is fed to the extruder at a defined feed rate,and the defined feed rate and the defined distance are selected so thatthe gel dispersion solidifies to form a tubular tissue scaffoldcomprising a biopolymer having fibrils in the tube wall that are alignedin a helical pattern around the longitudinal axis of the tube and wherethe pitch of the helical pattern on the luminal surface of the tube isdifferent from the pitch of the helical pattern on the outside surfaceof the tube.
 16. The method according to claim 15, wherein the outsidediameter of the annular-shaped exit port is about 5 mm and the width ofthe annular space is about 0.5 mm, and wherein the defined feed rate issufficient to provide a tube extrusion rate of about 150 cm/min. and thedefined distance is between about 10 cm and about 20 cm.
 17. The methodaccording to claim 13, where the tubular tissue scaffold remains in thebath for about 15 min.
 18. The method according to claim 3, wherein thestep of causing the gel to solidify includes immersing the tube in anaqueous solution containing 0.3% sodium bicarbonate, percent by weight.19. The method according to claim 3, further comprising sterilizing thetubular tissue scaffold with gamma or UV radiation.
 20. The methodaccording to claim 3, wherein the counter-rotating cone extrudercomprises an external rotating member having a cone-shaped cavitytherethrough and an internal rotating cone which fits within thecone-shaped cavity of the external rotating member, and wherein theexternal rotating member is driven to rotate in one direction at a speedof from about 1 to about 1800 rpm and the internal rotating cone isdriven to rotate in the opposite direction at a speed of from about 1 toabout 1800 rpm.
 21. The method according to claim 20, wherein theexternal rotating member is driven to rotate in one direction at a speedof from about 150 to about 900 rpm and the internal rotating cone isdriven to rotate in the opposite direction at a speed of from about 150to about 900 rpm.